The present invention relates to the magnetic resonance arts. It finds particular application to human magnetic resonance imaging in which a radio frequency magnetic resonance imaging coil is tuned to the resonance frequencies of phosphorous and hydrogen (or other dipoles of interest) and will be described with particular reference thereto. The coil may be used in transmit and receive modes, in transmit only modes, in receive only modes and may be used in conjunction with local gradients for high resolution imaging as well as spectroscopy. It is to be appreciated, however, that the invention will also find application in animal studies, non-human studies, high resolution in-vitro cell culture and profusion studies, industrial applications, security inspections and the like, as well as in other types of magnetic resonance imaging systems, magnetic resonance spectroscopy systems, and the like.
In magnetic resonance imaging applications, radio-frequency (RF) coils are used both to transmit RF pulses into and receive induced nuclear magnetic resonance (NMR) signals from a bore of a magnetic resonance imager. Typically, magnetic resonance imagers also include a series of annular resistive or superconducting magnets. Vacuum dewars in superconducting magnets and housing structures of resistive magnets define a central, longitudinal bore within which the subject is received. Commonly, a series of gradient magnetic field coils are mounted to a cylindrical dielectric former(s) for mounting in the magnet bore, reducing the patient receiving diameter. A whole body RF coil is mounted on another dielectric former for mounting in the interior bore of the gradient coil dielectric former. The gradient and RF coils, RF shield, and the formers are potted into a unitary structure, further reducing the patient receiving diameter.
Maintaining a large bore or patient aperture is advantageous. A large patient receiving bore not only accommodates large patients and provides a less claustrophobic environment, but it also allows imaging of portions of the subject further from the center of the bore. For example, shoulder imaging requires the patient's shoulders to be displaced radially inward from the RF coil.
The diameter reductions in the bore become more critical when self-shielded gradient coils are used. With self-shielded gradient coils, there are two sets of gradient coils disposed in a spaced relationship. The pair of gradient coil sets produce magnetic fields which (1) sum within the bore to create the desired magnetic field gradients and (2) subtract outside the bore. The subtraction zeroes or minimizes the external field to inhibit magnetic field gradient pulses from inducing eddy currents in the main magnet and associated structures. To achieve this shielding effect efficiently, a significant minimum spacing between the primary and secondary gradient coils is required. Analogously, an RF shield is advantageously disposed between the RF coil and the gradient coils to prevent the RF pulses from inducing eddy currents in the gradient coils. Again, a significant, minimum spacing between the RF coils and the RF shield is required.
Typically, a whole-body RF coil consists of a copper pattern etched on a circuit board with an FR-4 dielectric or other dielectric form. One or more of these circuit boards are wrapped around and fastened to the outer diameter of a filament wound tube or other cylindrical or elliptical dielectric form. Electrical components, such as capacitors, inductors, and diodes are soldered on to the circuit board to resonate the coil at the desired frequency. The tube has a diameter of 55 to 60 cm or about 20 mm or more less than the interior diameter of the gradient coil in which it is inserted, depending on the size of the patient. A tube of 55 to 60 cm is large enough to allow patients of up to approximately 180 kg to fit inside. This tube is inserted inside the gradient coil assembly and within the RF shield of the coil assembly. Feet, skid bars, or other support structures are used to position the coil concentric with the gradient tube and, thus, the RF shield.
The use of the filament wound tube has significant disadvantages. Importantly, it decreases the amount of radial space available to the patient. Further, an appropriately constructed and machined tube adds significantly to the cost of the body coil. In addition, when a filament wound tube is used, the construction and assembly of additional parts, such as feet and skid bars, further increase the total cost. Further, when a filament wound tube is used with an integral RF shield, it is difficult to ensure concentricity of the RF coil with the RF shield. A departure from concentricity adversely affects the performance of the RF coil by decreasing the coil's quadrature isolation (and thus signal-to-noise ratio) and uniformity. This becomes even more critical when the RF coil is brought closer to the RF shield.
The frequency at which the coil needs to operate is dependent on the main magnet's field strength B.sub.0 and the gyromagnetic ratio .gamma. of the dipole(s) of interest. In a 1.5 Tesla magnetic field, hydrogen dipoles .sup.1 H have a resonance frequency of about 64 MHz. There are typically other dipoles in the examination region with markedly different resonance frequencies, e.g., phosphorous .sup.31 P with a resonance frequency about 24 Mhz, or flourine .sup.19 F with a resonance frequency of 60 MHz. The radio frequency coil is commonly tuned to the magnetic resonance frequency of the selected dipole of interest.
When performing magnetic resonance spectroscopy upon a subject, it is desirable to switch quickly between proton imaging and spectroscopy imaging. Coils which can operate at more than one resonant frequency have traditionally been of the following types: doubly-tuned; electrically-switched; and, inserted-sleeve.
A doubly-tuned RF coil is simultaneously resonant at two separate frequencies. The coil uses both a high pass and a low pass circuit. At the lower frequency, the low pass circuit resonates, while at the higher frequency, the high pass circuit resonates. Alternatively, doubly-tuned RF coils are constructed by placing a tank circuit in series with resonant capacitors. The extra capacitance is switched in at one frequency. At the second resonant frequency, the tank circuit blocks current from flowing in these extra capacitors. Although some nuclei, such as phosphorous, lend themselves well to doubly-tuned coils of either type, others, such as fluorine, do not. This is because the imaging frequency of fluorine is very close to the imaging frequency of hydrogen. With such close frequencies, i.e., separated by less than an octave, the coil has difficulty resonating separately at both frequencies. One of the two resonant frequencies operates in a non-optimum mode, thus sacrificing the performance of the coil at that frequency. In addition, it is difficult if not impossible to tune a birdcage coil to more than two separate frequencies.
Although other methods for double-tuning a birdcage coil have been described, such as Peter Joseph et al., "A Technique for Double Resonant Operation of Birdcage Imaging Coils," IEEE Trans. Medical Imaging, vol. 8, No. 3, September 1989, and U.S. Pat. No. 5,212,450 by Murphy-Boesch et al., each method has its distinct disadvantages. The method of Joseph introduces field inhomogeneities while the method of Murphy-Boesch et al. increases the complexity of the design and construction by requiring four rings. In general, doubly-tuned methods do not allow for quadrature imaging. An exception is the four ring design of Murphy-Boesch et al.
Electrically-switched RF coils employ PIN diodes to switch between two sets of capacitors in order to adjust the tuning. Such coils have the disadvantage of working only for receive coils and thus are not applicable for whole body coils. Further, the high voltage and currents generated during transmit pulses can short circuit the diodes and render the coil ineffective.
Inserted-sleeve coils use a dielectric tube inside the RF coil to create an additional separate resonance frequency. See U.K. Patent Application No. 89-25919.6 by Schnur. On its outer diameter, the dielectric tube has metallic strips which add capacitance in parallel to the existing coil capacitors. Another type of inserted-sleeve coil uses two separate RF shields. See D. Lu and P. M. Joseph, "A Technique of Double Resonant Operation of F-19 and H-1 Quadrature Birdcage Coils," 1990 SMRM Book of Abstracts, p. 531. The two shields have such diameter that each one causes the RF coil to resonate at the desired frequency. These RF coils have two disadvantages. First, it is difficult to insert a sleeve without repositioning the patient. Second and more importantly, these methods only work when the two desired resonances are very close in frequency. It should also be noted that RF coils using dielectric tubes inside the coils have been found to be inherently unstable over extended periods of time (i.e., the resonant frequency of the coil tends to vary with time and temperature).
The present invention contemplates a new and improved method and apparatus for positioning a whole-body RF coil and mechanically switching its resonant frequencies within a bore of an MRI apparatus which overcomes the above-mentioned problems and others.